The present invention relates to the magnetic resonance arts. It finds particular application in conjunction with a radio frequency magnetic resonance imaging coil which is tuned to the resonance frequencies of hydrogen (or other dipoles of interest) for anatomical, angiographic, functional and other medical imaging of humans and will be described with particular reference thereto. It is to be appreciated, however, that the invention will also find application in animal studies, other non-human studies, spectroscopy, phased-array coil techniques, and the like.
Conventionally, magnetic resonance imaging systems generate a strong, uniform static magnetic field in a free space or bore of a magnet. This main magnetic field polarizes the nuclear spin system of an object in the bore to be imaged. The polarized object then possesses a macroscopic magnetic moment vector pointing in the direction of the main magnetic field. In a superconducting main annular magnet assembly, the static magnetic field B.sub.0 is generated along a longitudinal or z-axis of the cylindrical bore.
To generate a magnetic resonance signal, the polarized spin system is excited by applying a magnetic resonance signal or radio frequency field B.sub.1, perpendicular to the z-axis. The frequency of the magnetic resonance signal is proportional to the gyromagnetic ratio .gamma. of the dipole(s) of interest. The radio frequency coil is commonly tuned to the magnetic resonance frequency of the selected dipole of interest, e.g., 64 MHZ for hydrogen dipoles .sup.1 H in a 1.5 Tesla magnetic field.
Typically, a radio frequency coil for generating the magnetic resonance signal is mounted inside the bore surrounding the sample or patient. In a transmission mode, the radio frequency coil is pulsed to tip the magnetization of the polarized sample away from the z-axis. As the magnetization precesses around the z-axis back toward alignment, the precessing magnetic moment generates a magnetic resonance signal which is received by the radio frequency coil in a reception mode.
For imaging, a magnetic field gradient coil is pulsed for spatially encoding the magnetization of the sample. Typically, the gradient magnetic field pulses include gradient pulses pointing in the z-direction but changing in magnitude linearly in the x, y, and z-directions, generally denoted G.sub.x, G.sub.y, and G.sub.z, respectively. The gradient magnetic fields are typically produced by a gradient coil which is located inside the bore of the magnet and outside of the radio frequency coil.
Conventionally, when imaging the torso, a whole body radio frequency coil is used in both transmit and receive modes. By distinction, when imaging the head, neck, shoulders, or an extremity, the whole body coil is often used in the transmission mode to generate the uniform excitation field B.sub.1 and a local coil is used in the receive mode. Placing the local coil close to the imaged region improves the signal-to-noise ratio and the resolution. In some procedures, local coils are used for both transmission and reception. One drawback to local coils it that they tend to be relatively small.
One type of local frequency coil is known as the "birdcage" coil. See, for example, U.S. Pat. No. 4,692,705 of Hayes. Typically, a birdcage coil has a pair of circular end rings which are bridged by a plurality of equi-spaced straight segments or legs. In a primary mode, currents in the rings and legs are sinusoidally distributed which results in improved homogeneity along the axis of the coil. Homogeneity along the axis perpendicular to the coil axis can be improved to a certain extent by increasing the number of legs in the coil. Typically, a symmetric birdcage coil has eight-fold symmetry. Such a symmetric birdcage coil with N legs exhibits N/2 mode pairs. The first (N/2)-1 mode pairs are degenerate, while the last mode pair is non-degenerate. The primary mode of such an eight-fold symmetric birdcage coil has two linear modes which are orthogonal to each other. The signals from these two orthogonal or quadrature modes, when combined, increase signal-to-noise ratio on the order of 40%. The simplest driven current pattern or standing waves are defined by superpositions of degenerate eigenfunctions. For a low-pass birdcage of n meshes driven at its lowest non-zero eigenfrequency, the current in the n-th mesh is given by sin(2.pi.n/N+.phi.). The phase angle .phi. determines the polarization plane of the resulting B.sub.1 radio frequency field and can be varied continuously by suitable application of driving voltages. The alignment and isolation of the two linear modes of a birdcage coil are susceptible to sample geometry. That is, the sample dominates the mode alignment and isolation between the two linear modes.
A typical 16-legged, high-pass birdcage coil has a diameter and length of about 30 centimeters. Capacitors interrupt the end rings between adjacent legs for a total of 32 such capacitors which are evenly distributed through the two end rings. Such a coil exhibits 7 degenerate modes and one non-degenerate mode. The principal or k=1 mode is tuned to approximately 63.72 MHZ which is the magnetic resonance frequency of protons in a 1.5 T static magnetic field. The birdcage coil uses a four-port electric current feed.
Generally, when used with annular superconducting magnets, the local RF head coil is oriented such that the coil axis is parallel to the magnetic axis. This enables the patient to access the coil volume easily. Further, this orientation takes advantage of the quadrature aspect of the local head coil with respect to the orientation of the main magnetic B.sub.0 field. With this arrangement, the legs are parallel to the horizontal magnetic axis.
One problem with local RF head coils is their claustrophobic effect on patients. Many pediatric and adult patients already experience claustrophobic reactions when placed inside the horizontal bore of a superconducting magnet. Placement of a close-fitting head coil having anterior legs which obstruct the direct view of the patients further adds to their discomfort. The discomfort is somewhat reduced by illumination inside the magnet bore, air flow and the use of reflective mirrors. However, these remedies are not enough to eliminate the claustrophobia problems.
Another problem with conventional head coils is that the design limits access to the patient. For example, it is often desirable to perform interventional medicine or use life-support devices, such as ventilator tubes, while imaging a patient. However, the proximity of the axial segments to one another and to the head of the patient impair such practices.
The present invention contemplates a new and improved radio frequency coil design which overcomes the above-referenced problems and others.